Optical coherence tomography (“OCT”) is an imaging technique that can measure an interference between a reference beam of light and a beam reflected back from a sample. A detailed system description of traditional time-domain OCT is described in Huang et al., “Optical Coherence Tomography,” Science 254, 1178 (1991). Optical frequency domain imaging (“OFDI”) techniques, which can be also known as swept source or Fourier-domain optical coherence tomography (OCT) techniques, can be OCT procedures which generally use swept laser sources. For example, an optical beam is focused into a tissue, and the echo time delay and amplitude of light reflected from tissue microstructure at different depths are determined by detecting spectrally resolved interference between the tissue sample and a reference as the source laser wavelength is rapidly and repeatedly swept. A Fourier transform of the signal generally forms an image data along the axial line (e.g., an A-line). A-lines are continuously acquired as the imaging beam is laterally scanned across the tissue in one or two directions that are orthogonal to the axial line.
The resulting two or three-dimensional data sets can be rendered and viewed in arbitrary orientations for gross screening, and individual high-resolution cross-sections can be displayed at specific locations of interest. This exemplary procedure allows clinicians to view microscopic internal structures of tissue in a living patient, facilitating or enabling a wide range of clinical applications from disease research and diagnosis to intraoperative tissue characterization and image-guided therapy. Exemplary detailed system descriptions for spectral-domain OCT and Optical Frequency Domain Interferometry are described in International Patent Application PCT/US03/02349 and U.S. Patent Application Ser. No. 60/514,769, respectively.
A contrast mechanism in the OFDI techniques can generally be an optical back reflection originating from spatial reflective-index variation in a sample or tissue. The result can be a so-called an “intensity image” that may indicate the anatomical structure of tissue up to a few millimeters in depth with spatial resolution ranging typically from about 2 to 20 μm. While the intensity image can provide a significant amount of morphological information, birefringence in tissues may offer another contrast useful in several applications such as quantifying the collagen content in tissue and evaluating disease involving the birefringence change in tissue. Polarization-sensitive OCT can provide an additional contrast by observing changes in the polarization state of reflected light. The first fiber-based implementation of polarization-sensitive time-domain OCT is described in Saxer et al., “High-speed fiber-based polarization-sensitive optical coherence tomography of in vivo human skin,” Opt. Lett. 25, 1355 (2000).
In polarization-sensitive time-domain OCT techniques, a simultaneous detection of interference fringes in two orthogonal polarization channels can facilitate a complete characterization of a reflected polarization state, as described in J. F. de Boer et al., “Determination of the depth-resolved Stokes parameters of light backscattered from turbid media by use of polarization-sensitive optical coherence tomography,” Opt. Lett. 24, 300 (1999). There can be two non-depolarizing polarization parameters: birefringence, characterized by a degree of phase retardation and an optic axis orientation, and diattenuation, which can be related to dichroism and characterized by an amount and an optic axis orientation. Together, these optical properties may be described by, e.g., the 7 independent parameters in the complex 2×2 Jones matrix.
The polarization state reflected from the sample can be compared to the state incident on the sample quite easily in a bulk optic system, as the polarization state incident on the sample can be controlled and fixed. However, an optical fiber may have a significant disadvantage in that a propagation through optical fiber can alter the polarization state of light. In this case, the polarization state of light incident on the sample may not be easily controlled or determined. In addition, the polarization state reflected from the sample may not be necessarily the same as the polarization state received at the detectors. Assuming negligible diattenuation, or polarization-dependent loss, optical fiber changes the polarization states of light passing through such fiber in such a manner as to preserve the relative orientation between states. The overall effect of propagation through optical fiber and non-diattenuating fiber components can be similar to an overall coordinate transformation or some arbitrary rotation. In other words, the relative orientation of polarization states at all points throughout propagation can be preserved, as described in U.S. Pat. No. 6,208,415.
There have been a number of approaches that can take advantage to determine the polarization properties of a biological sample imaged with polarization-sensitive OCT. Such approaches have suffered from some disadvantage, however.
For example, a vector-based method has been used to characterize birefringence and optic axis orientation only by analyzing rotations of polarization states reflected from the surface and from some depth for two incident polarization states perpendicular in a Poincaré sphere representation as described in the Saxer Publication, J. F. de Boer et al., “Determination of the depth-resolved Stokes parameters of light backscattered from turbid media by use of polarization-sensitive optical coherence tomography,” Opt. Lett. 24, 300 (1999), and B. H. Park et al., “In vivo burn depth determination by high-speed fiber-based polarization sensitive optical coherence tomography,” J. Biomed. Opt. 6, 474 (2001).
Mueller matrix based methods are capable of determining birefringence, diattenuation, and optic axis orientation as described in S. L. Jiao et al., “Two-dimensional depth-resolved Mueller matrix of biological tissue measured with double-beam polarization-sensitive optical coherence tomography,” Opt. Lett. 27, 101 (2002), S. Jiao et al., “Optical-fiber-based Mueller optical coherence tomography,” Opt. Lett. 28, 1206 (2003), and S. L. Jiao et al., “Depth-resolved two-dimensional Stokes vectors of backscattered light and Mueller matrices of biological tissue measured with optical coherence tomography,” Appl. Opt. 39, 6318 (2000). These typically utilize a multitude of measurements using a combination of incident states and detector settings and limits their practical use for in vivo imaging.
Jones matrix based approaches have also been used to characterize the non-depolarizing polarization properties of a sample as described in S. Jiao et al., “Optical-fiber-based Mueller optical coherence tomography,” Opt. Lett. 28, 1206 (2003) and S. L. Jiao and L. V. Wang, “Jones-matrix imaging of biological tissues with quadruple-channel optical coherence tomography,” J. Biomed. Opt. 7, 350 (2002). The description of these approaches has limited a use of optical fiber and fiber components such as circulators and fiber splitters such that these components must be traversed in a round-trip fashion and assumes that sample birefringence and diattenuation share a common optic axis. These approaches can use a multitude of measurements using a combination of incident states and detector settings and limits their practical use for in vivo imaging.
Generally, in nearly all of polarization sensitive time domain, Spectral Domain OCT, or OFDI systems, the polarization properties can be measured using different incident polarization states on the sample in a serial manner, i.e. the incident polarization state incident on the sample was modulated as a function of time.
Exemplary system and method for obtaining polarization sensitive information is described in U.S. Pat. No. 6,208,415. Exemplary OFDI techniques and systems are described in International Application PCT/US04/029148. Method and system to determine polarization properties of tissue is described in International Application PCT/US05/039374.
For example, rotation and/or translation of an imaging catheter can give rise to rotations of a polarization mode along single mode optic fibers used to guide electromagnetic radiation (e.g., light). This can occurs within a single rotation of the imaging catheter, and can therefor therefore impact the uniformity of the input polarization state within a single imaging frame, making it impossible to reconstruct accurate information about the orientation of birefringent tissue that would otherwise be available.
Polarization sensitive optical coherence (“PS-OCT”) technology imaging in fiber-based systems typically calculate the magnitude of birefringence while disregarding the vectorial aspect of the acquired data. This is because in single mode fibers random variations in the shape of the fiber core, either innate or due to external stress/strain, result in additional birefringence that renders absolute measurements of sample birefringence impossible. Instead, differential measurements of sample birefringence can be required starting at the sample surface, which serves as a calibration reflector. As a result, the magnitude of the birefringence can be recovered, although the orientation may be lost. Attempts to recover this information have heretofore involved using polarization—maintaining fibers [see, e.g., Ref 1] or placing bulk waveplates in the field of view between the imaging lens and sample to calibrate the orientation of the probing light [see, e.g., Ref. 2]. While these approaches can be successful when imaging via a bench top microscope configuration, neither is applicable when performing catheter imaging in in vivo applications. In addition, catheter imaging can be further complicated because the rotation of the catheter affects the measured orientation over the course of a single frame, so that from A-line to A-line the measured state is spuriously altered.
In systems that operate as described above, polarizing beam splitters send orthogonal components of the interference signal to two detectors, allowing the Jones vector to be determined. The signal is related to the source field by
      E    f    =            (                                                                  E                fx                            ⁢                              ⅇ                                  ⅈ                  ⁢                                                                          ⁢                                      φ                    fx                                                                                                                                          E                fy                            ⁢                              ⅇ                                  ⅈ                  ⁢                                                                          ⁢                                      φ                    fy                                                                                          )        =                  ⅇ                  ⅈ          ⁢                                          ⁢          θ                    ⁢              J        out            ⁢              J        S            ⁢                        J          in                ⁡                  (                                                                                          E                                          θ                      ⁢                                                                                          ⁢                      x                                                        ⁢                                      ⅇ                                          ⅈ                      ⁢                                                                                          ⁢                                              φ                                                  θ                          ⁢                                                                                                          ⁢                          x                                                                                                                                                                                                              E                                          θ                      ⁢                                                                                          ⁢                      y                                                        ⁢                                      ⅇ                                          ⅈ                      ⁢                                                                                          ⁢                                              φ                                                  θ                          ⁢                                                                                                          ⁢                          y                                                                                                                                                  )                    where the J′S represent the Jones matrices of the output path, the sample, and the input path, respectively, and θ is the common phase. At a reference reflecting surface is JS=1. Birefringence measurements at greater depths in the sample arm can then be related to this calibration measurement to yield the Jones matrix JA=JoutJSJ−1out. Assuming no diattenuation exists in the system, Jout represents a unitary transformation and therefore the relative retardance and optic axis of the sample may be extracted straightforwardly. The absolute orientation in the laboratory frame is, however, completely obliterated by the effect Jout has on the measurements.
In the processing of the raw data, the complex components of the Jones vector measured by such system can be converted into the Stokes parameters Q, U, and V, with Q and U representing the basis vectors of linearly polarized states and V that of circularly polarized states. Since all sample data are acquired in the reflection configuration and therefore involve round—trip measurements, circular birefringence cannot be detected due to the round—□trip cancellation of any circularly birefringent signal. The physical orientations corresponding to the remaining Stokes parameters Q and U are depicted in FIG. 6.
The technique used to extract the principal axis has been described in a number of publications [see, e.g., Refs. 3-5]. For example, it is possible to alternate between two input polarization states that are orthogonal on the Poincare sphere. Switching between the two input states occurs at a frequency that is half that at which the system acquires A-lines, so that each successive pair of A-lines together form a single birefringence measurement at a (roughly) identical location in the sample. The relative axis can then be determined by calculating the pair of difference Stokes vectors along the axial profile to find how the polarization states were rotated and taking the cross product in order to identify the direction vector that gave rise to the rotations. This can correspond to the optic axis of the tissue at that depth in the sample.
Although it is well-established that airway smooth muscle (ASM) abnormality is a primary pathophysiologic mechanism in asthma, clinical asthma research has predominately neglected the contribution of ASM in favor of focusing on other factors such as airway inflammation. This can be largely due to the lack of means for adequately assessing ASM in vivo. Asthma is a widespread problem affecting hundreds of millions of people, approximately 10% of which are estimated to suffer from uncontrolled or poorly controlled symptoms. The difficulty in providing an asthma patient with the proper treatment can be in part due to a lack of understanding with regards to the pathophysiology of asthma, and the challenges currently faced in gaining such knowledge. While it is believed that ASM proliferation is intimately connected to the airway hyper-responsiveness experienced by sufferers of asthma [see, e.g., Refs. 8-10], the full extent of the connection between asthma and ASM abnormality has proved difficult to assess, largely because there has existed no imaging technology with the spatial and temporal resolution necessary to visualize ASM distribution and dynamics in vivo. The situation is further complicated by the fact that, as we now know, asthma itself is not a single disease but rather a collection of phenotypes [see, e.g., Refs. 11 and 12].
Among currently available imaging modalities, optical coherence tomography (OCT) [see, e.g., Ref. 13] offers what is arguably the most promising balance of technological features for undertaking fully non-invasive, in vivo studying of ASM and asthma. However, structural OCT still lacks the contrast necessary to discriminate ASM bands on a level approaching that of histological staining. One potential approach to overcoming this is to take advantage of the form birefringence of ASM fibers by utilizing polarization sensitive imaging. Form birefringence is the quantifiable anisotropy of ordered material structure manifested in the differential propagation time of orthogonal polarization states of light transmitted through the structure. Polarization sensitive OCT (PS-OCT) introduces a way of obtaining information about this anisotropy, bringing additional contrast to standard OCT images [see, e.g., Refs. 14-17].
Accordingly, there may be a need to address and/or overcome at least some of the issues of deficiencies described herein above.